The present invention relates to the art of nuclear imaging. It finds particular application in relation to gamma cameras having a detector which utilizes a scintillating crystal.
In nuclear imaging, a radiopharmaceutical containing a radionuclide such as .sup.99m Tc or .sup.201 Tl is introduced into the body of a patient. As the radiopharmaceutical decays, gamma rays are generated. These gamma rays are detected and used to construct a clinically useful image.
The gamma rays are detected using one or more detectors. These detectors ordinarily include a scintillator crystal which emits photons or light energy in response to incident radiation such as a gamma ray or other high energy photon. An array of photomultiplier tubes (PMTs) is used to detect the light emitted by the scintillator crystal. The signals generated by the PMTs are in turn used to determine the location and energy of the detected event. This information is used to produce an image indicative of the patient's anatomy.
A sheet of optical glass has been placed adjacent the scintillator crystal on the side facing the source of radiation (i.e. on the side nearer the imaging region of the gamma camera). The sheet of optical glass has been bonded to the scintillator using a silicon optical adhesive in a procedure performed by the Bicron Technology business unit of St. Gobain/Norton Industrial Ceramics Corporation, located in Newbury, Ohio. The bond allows the crystal and the optical glass to be mechanically joined without adversely affecting the path of photons through the interface between the glass and crystal. Hence, one function of the optical glass is to provide structural support to the scintillator crystal.
Conventionally, the scintillating crystals used in gamma cameras have had a thickness of 0.375 inches (0.9525 cm), and the optical glass has had a thickness of 0.675 inches (1.7145 cm). The bonding material has a nominal thickness of approximately 0.030 inches (0.0762 cm) such that the total thickness of the crystal-glass structure is approximately 1.030 inches (2.62 cm).
A honeycomb of a nickel alloy commonly known as mu-metal has, together with the PMTs, been pressed against the rear surface of the optical glass. The mu-metal structure defines a plurality of hexagonal apertures into which the PMTs have been placed. The mu-metal structure surrounding each of the PMTs reduces the effects of the earth's magnetic field as the detector is moved about the patient.
A number of factors make it desirable to vary the relative thicknesses of the optical glass and the scintillator crystal. One example arises in positron emission tomography (PET), a branch of nuclear medicine in which a positron emitting radiopharmaceutical such as .sup.18 F-fluorodeoxyglucose (FDG) is introduced into the body of a patient. Each emitted positron reacts with an electron in what is known as an annihilation event, thereby generating a pair of gamma rays which are emitted in directions approximately 180 degrees apart, i.e. in opposite directions. The gamma rays produced by a positron annihilation are characterized by a photopeak at 511 keV, as compared to a 140 keV photopeak for .sup.99m Tc.
A pair of detectors registers the positions and energy of the respective gamma rays, thereby providing information as to the position of the annihilation event and hence the positron source. Because the gamma rays travel in opposite directions, the positron annihilation is said to have occurred along a line of coincident connecting the detected gamma rays. A number of such events are collected and used to reconstruct a clinically useful image.
Sensitivity and resolution are important gamma camera characteristics. A higher sensitivity permits the use of smaller doses of radiopharmaceutical. For a given amount of incident radiation, the gamma camera detects a larger number of events and thereby produces images having greater diagnostic utility.
One factor which affects the sensitivity and resolution of a gamma camera is the efficiency of its scintillating crystal. In fact, many of the incident gamma rays pass through the crystal without any interaction and are thus not detected by the gamma camera. The efficiency of the crystal is also function of the energy of the gamma radiation. For example, conventional NaI(Tl) crystals have a lower efficiency at energies associated with PET imaging than at energies associated with more conventional nuclear imaging.
In order to improve gamma camera performance, it is becoming increasingly important to increase the efficiency of the scintillating crystal. For example, nuclear cameras are increasingly being used to perform positron annihilation imaging, with its relatively higher energies. It is also desirable to increase the efficiency of the scintillating crystal at lower energies.
One technique for improving the efficiency of the scintillator crystal is to increase its thickness. Gamma rays passing through the crystal are thus more likely to interact with the crystal and thus produce a flash of light detectable by the PMTs. There are, however, a variety of practical considerations which make it difficult to simply increase the thickness of the crystal.
Among these considerations is the fact that optical design constraints limit the overall thickness of the detector. Moreover, retrofitting thicker crystals on existing cameras and camera designs is facilitated if the improved crystal structure fits in the pre-existing support structure. Further, adequate structural support for the relatively fragile crystal must still be provided. Thus, a technique which facilitates the use of a thicker scintillating crystal while providing adequate structural support in a size efficient package is needed.